The present invention relates to compositions and methods for treating disorders associated with tissue damage, loss, or malformation.
Tissue engineering, i.e., the generation of new living tissues in vitro, is widely used to replace diseased, traumatized or other unhealthy tissues. The classic tissue engineering approach utilizes living cells and a basic scaffold for cell culture (Langer and Vacanti, 1993; Nerem and Seliktar, 2001). Thus, the scaffold structure attempts to mimic the natural structure of the tissue it is replacing and to provide a temporary functional support for the cells (Griffith L G, 2002).
Tissue engineering scaffolds are fabricated from either biological materials or synthetic polymers. Synthetic polymers such as polyethylene glycol (PEG), Hydroxyapatite/polycaprolactone (HA/PLC), polyglycolic acid (PGA), Poly-L-lactic acid (PLLA), Polymethyl methacrylate (PMMA), polyhydroxyalkanoate (PHA), poly-4-hydroxybutyrate (P4HB), polypropylene fumarate (PPF), polyethylene glycol-dimethacrylate (PEG-DMA), beta-tricalcium phosphate (beta-TCP) and nonbiodegradable polytetrafluoroethylene (PTFE) provide precise control over the physical properties of the material (Drury and Mooney, 2003).
Common scaffold fabrication methods are based on foams of synthetic polymers. However, cell migration into the depth of synthetic scaffolds is limited by the lack of oxygen and nutrient supply. To overcome such limitations, new approaches utilizing solid freeform fabrications and internal vascular architecture have been developed (Reviewed in Sachlos E and Czernuszka J T, 2003; Eur. Cell Mater. 5: 29-39). Likewise, freeze-drying methods are also employed to create unique three-dimensional architectures with distinct porosity and permeability. However, creating pores into these materials is an aggressive procedure, often involving the use of toxic conditions which eliminate the possibility of pre-casting tissue constructs with living cells. Therefore, many of the prefabricated materials are subject to uneven cell seeding and inhomogeneous populations of cells within the constructs. Furthermore, the materials are generally degraded unevenly during the tissue cultivation process, creating a highly anisotropic tissue with altered growth kinetics.
Scaffolds made of PEG are highly biocompatible (Merrill and Salzman, 1983) and exhibit versatile physical characteristics based on their weight percent, molecular chain length, and cross-linking density (Temenoff J S et al., 2002). In addition, PEG hydrogels are capable of a controlled liquid-to-solid transition (gelation) in the presence of cell suspension (Elbert and Hubbell, 2001). Moreover, the PEG gelation (i.e., PEGylation) reaction can be carried out under non-toxic conditions in the presence of a photoinitiator (Elisseeff J et al., 2000; Nguyen and West, 2002) or by mixing a two-part reactive solution of functionalized PEG and cross-linking constituents (Lutolf and Hubbell, 2003).
However, while the abovementioned synthetic polymers enable precise control over the scaffold material, they often provide inadequate biological information for cell culture. As a result, these materials are unsuitable for long-term tissue culture or in vivo tissue regeneration.
On the other hand, naturally occurring scaffolds such as collagen, fibrin, alginate, hyaluronic acid, gelatin, and bacterial cellulose (BC) provide bio-functional signals and exhibit various cellular interactions. For example, fibrin, a natural substrate of tissue remodeling (Herrick S., et al., 1999), contains several cell-signaling domains such as a protease degradation substrate (Werb Z, 1999) and cell-adhesion domains (Herrick S., 1999). However, because such biological materials exhibit multiple inherent signals (e.g., regulation of cell adhesion, proliferation, cellular phenotype, matrix production and enzyme activity), their use as scaffolds in tissue regeneration often results in abnormal regulation of cellular events (Hubbell, 2003). Furthermore, the natural scaffolds are often much weaker after reconstitution as compared to the strength of the original biological material, and little control can be exercised to improve their physical properties.
Another drawback of natural scaffolds (e.g., collagen and fibrin) for tissue engineering is the limited control over the physical properties of the polymeric network. For example, reconstituted collagen undergoes fibrilogenesis and self-assembly to form an interpenetrating network of nano-scale fibrils that loosely associate together by non-specific interactions such as hydrogen bonding. In comparison to the highly organized and enzymatically cross-linked collagen fibers of the normal tissue structure, the interpenetrating network of fibrils exhibit poor physical strength and super-physiological tissue porosity. Moreover, the specific conformation of fibrils combined with the open pore structure of the interpenetrating network leaves the protein backbone easily accessibly and susceptible to freely diffusing proteases from the surrounding host tissue or cell culture system. This often results in uncontrolled and premature deterioration of the scaffold in the presence of cell-secreted proteases. The discrepancies in structure and function of reconstituted protein hydrogels compared to natural urges the development of biomimetic scaffold systems for implementation in many practical tissue engineering applications.
To date a number of techniques have been developed for the modification and improvement of the physicochemical properties of reconstituted protein hydrogels which prevent them from rapid degradation. Collagen and fibrin hydrogels can be processed by freeze-drying the construct to increase the tensile strength and modulus of the protein network. However, freeze-drying necessitates a pre-fabrication freezing step which eliminates the possibility for gelation of the polymer in the presence of cells and also eliminates the benefits of in-situ polymerization. The freeze-drying process also affects the molecular architecture of the polymer network, altering the nano-fiber mesh and turning it into a macro-porous sponge structure. Other techniques for improving the physical properties of natural hydrogels while maintaining the nano-fiber structure have been proposed based on covalent cross-links, including the use of aldehydes, carbodiimides [Park S N, Park J C, Kim H O, Song M J, Suh H. Characterization of porous collagen/hyaluronic acid scaffold modified by 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide cross-linking. Biomaterials. 2002 February; 23(4):1205-12], and N-hydroxysuccinimides (NHS) in the presence of amino acids [Ma L, Gao C, Mao Z, Zhou J, Shen J. Enhanced biological stability of collagen porous scaffolds by using amino acids as novel cross-linking bridges. Biomaterials. 2004 July; 25(15):2997-3004; Ma L, Gao C, Mao Z, Zhou J, Shen J. Biodegradability and cell-mediated contraction of porous collagen scaffolds: the effect of lysine as a novel crosslinking bridge. J Biomed Mater Res A. 2004 Nov. 1; 71(2):334-42]. All the cross-linking procedures offer some improvements of the physical stability of the scaffold, but do so by introducing a toxic manufacturing step which requires extensive washes and increases the likelihood that residual toxins in the scaffold will affect cellular activity.
The proteolytic degradation of protein scaffolds can also be delayed by protecting the protein backbone of the polymer network using covalent attachment of a shielding polymers such as poly(ethylene glycol) (PEG). For example, the modification of proteins by attachment of one or more PEG chains (PEGylation) has been applied very successfully to increasing the plasma half-life of therapeutic peptides or protein drugs [Veronese F M. Peptide and protein PEGylation: a review of problems and solutions. Biomaterials. 2001 March; 22(5):405-17]. Based on a similar rationale, PEGylation could be a good strategy for protein-based biomaterial design in as much as the PEG chains can slow down the enzymatic biodegradation of the PEGylated protein scaffold. At the same time, the PEG chains are non-toxic, non-immunogenic, highly water soluble, and are already approved by the FDA in a number of different clinical indications (Veronese, 2005). Common proteins used in scaffold design such as collagen and fibrin may be readily PEGylated using amine group modifications or thiol modifications of the protein backbone to yield a protein-polymer conjugate. The PEG shields the protein surface from degrading agents by steric hindrances without blocking all the natural biological function of the structural protein molecule (Veronese 2005).
Recently hybrid scaffolds have been developed. A hybrid scaffold material combines the structural characteristics of the synthetic material with the biofunctionality of natural material (Leach J B, et al., 2004; Leach and Schmidt, 2005). To this end, several methods of preparing scaffold with natural biofunctionality and physical properties of synthetic polymers have been proposed. Most of these “hybrid” approaches, however, fall short of producing a biomaterial with broad inherent biofunctionality and a wide range of physical properties; mainly because they employ only a single biofunctional element into the material design. For example, prior studies describe the preparation of scaffolds consisting of biodegradable elements grafted into the backbone of a synthetic hydrogel network. Hydrogels were prepared from synthetic PEG which was cross-linked with short oligopeptides containing enzymatic substrates capable of being proteolytically degraded by cell-secreted enzymes [Lutolf et al (2003); Gobin and West (2002)]. Furthermore, to increase the biofunctionality of such hydrogels, synthetic adhesion motifs such as the RGD sequences [Lutolf et al (2003)] or VEGF (Seliktar et al; 2004, Zisch A H, et al, 2003; FASEB J. 17: 2260-2. Epub 2003 Oct. 16) were grafted into the PEG backbone. However, the use of such scaffolds (in which PEG is the major component) was limited by the insufficient bio-feedback and/or long-term cellular responses which are essential for phenotypic stability.
Further attempts to increase the biofunctionality of the scaffolds included the manufacture of genetically-engineered protein-like precursors of 100 amino acids, which contain, among other things, several protease substrates and adhesion sites (Halstenberg et al. 2002; Biomacromolecules, 3: 710-23). However, the increased protein precursors size and the presence of thiol groups required for the PEGylation reaction complicated the purification and solubilization of the precursors during the scaffold manufacturing process. In addition, similar to the PEG-based biosynthetic materials, the genetically-engineered protein precursor scaffolds failed to provide sufficient biofunctionality to enable long-term stability.
The present inventor has previously uncovered that biosynthetic hybrid scaffolds composed of a fibrinogen backbone which is cross-linked with functional polyethylene glycol (PEG) side chains are excellent, biodegradable scaffolds which can be used for tissue regeneration applications.
While reducing the present invention to practice, the present inventor has uncovered that the above scaffold are subject to the proteolytic and hydrolytic activity of the cellular environment in the implantation site causing sustained release of PEGylated denatured fibrinogen degradation products. These PEGylated denatured fibrinogen degradation products have similar inductive properties of the natural fibrin degradation products with the added advantage of the PEG modification which provides protection from rapid clearance from the local implantation site and from the body.